Many crucial biological functions are mediated or accomplished by biomolecules and tissue structures that are intrinsically fluorescent. As a result, there is an opportunity to diagnose and study important biological events by measuring and localizing the spectra and tissue fluorescence emission. To investigate in vivo internal processes and structures in large organisms, such as human beings and agricultural animals, an endoscopic procedure which penetrates body cavities or even solid tissue may be required.
Endoscopy video imaging in body cavities ordinarily utilizes back-scattered white light applied through the endoscope to form a low-resolution color image of the internal surfaces of these cavities. Physicians often use the changes in shapes and changes in local apparent color (which are often due to changes in blood distribution) to recognize disease states, such as malignant tumors or inflammation. Unfortunately, these clues are frequently not sufficient, especially for detection of the early onset of disease. Diagnostic improvements have been made by quantitative measurements of the light scattering and of tissue fluorescence emission.
Ordinarily, the light required to excite the fluorescence of tissue is delivered through an optical fiber or fiber bundle that is inserted through a small tube built into the endoscopic pipe. Small optical fibers or fiber bundles for emission collection can be passed through the same tube, and/or the delivery fiber can be used for epi-collection. Some of the strongest tissue fluorescence usually seen in this procedure is due to NADH (nicotinamide adenine dinucleotide), collagen and aromatic amino acids, such as tryptophan. Their fluorescence is excited by absorption of ultraviolet light of about 250 to 450 nm wavelength corresponding to photon energies of around 2 to 4 eV or sometimes slightly longer wavelength visible light may be used for intrinsic emitters such as flavoproteins. It is also possible to excite and collect fluorescence emissions from added fluorescent dyes such as fluorescein to increase contrast and signal strength.
A first problem is that this light is strongly absorbed by hemoglobin and oxyhemoglobin in the blood so that penetration of the illumination into the tissue depends on their concentration and distribution.
A second problem is that the illumination exiting the optical fibers into tissue fans out at an included angle determined by the numerical aperture (NA) of the optical fiber. Small lenses can be used so that the light first converges to a focus but it then fans out beyond the focal plane. (Typically, the NA is about 0.2 and the included cone angle is ˜23° in air.) With single photon (i.e., linear) excitation this angular spreading is a problem, because equal total amounts of fluorescence are excited in every spherical section at each distance from the end of the fiber until attenuated by absorption and scattering. This effect is schematically illustrated in FIG. 1. Fluorescence excitation is similarly spread out. Scattering does not attenuate the fluorescence excitation but does distribute it even more broadly. Consequently, the volume observed is ill-defined with its practical limits depending also on blood distribution and light scattering. It should be noted that these problems tend to persist even if lenses focus the illumination and/or prisms and mirrors deflect the light for side viewing.
Femtosecond-pulse propagation through large-core microstructured fibers was investigated. Although these fibers are highly multimode, excitation of the fundamental mode is readily achieved, and coupling to higher-order modes is weak even when the fiber is bent or twisted. For prechirped input pulses with energies as large as 3 nanojoule (“nJ”), pulses as short as 140 femtoseconds (“fs”) were produced at the output of the fiber. Such a system could prove to be extremely useful for applications such as in vivo multiphoton microscopy and endoscopy that require delivery of femtosecond pulses and collection of fluorescence.
Femtosecond pulses generated in a diffraction-limited beam have become an important tool for many science and technology areas. For a number of applications, it is desirable to deliver these pulses via an optical fiber over a distance of a few meters to a specific location. It is commonly believed that one must use single-mode fibers (“SMFs”) to ensure the high spatial quality of the beam as well as to avoid additional temporal broadening as a result of intermodal dispersion. However, as a result of the very small (<5 μm) core size of SMFs, self-phase modulation broadens the pulse spectrum for pulse energies as low as a few picojoules, and the dispersion-induced temporal broadening cannot be readily compensated for by means of prechirping.
For applications such as multiphoton microscopy (Williams et al., Curr. Opin. Chem. Biol. 5:603 (2001)), fiber delivery of femtosecond pulses could improve the existing microscope design and, more importantly, lead to the development of miniaturized instrumentation for both basic research in biology and clinical applications such as nonlinear endoscopy. To preserve the pulse width after propagation of the pulse through the fiber, one must compensate for or minimize dispersive and nonlinear effects. Fiber dispersion can be compensated for by imparting a suitable frequency chirp on the input pulses. Propagation of negatively chirped pulses in a SMF was investigated (Myaing et al., Opt. Express 7:210 (2000)), and the width of the output pulses was found to scale sublinearly with pulse energy, but even for a pulse energy of 0.5 nJ the output pulse width exceeded 0.5 ps. Atherton and Reed (Atherton et al., Proc. SPIE 3269:22 (1998)) used a single-grating precompensator to deliver 100-fs pulses with energies as great as 0.7 nJ through a 3-m-long fiber. To minimize the nonlinearity, they stretched the pulse in the compensator by an amount that was more than the amount that was compensated for by the fiber, and an additional piece of positive dispersive glass was used to restore the pulse nearly to its initial duration. However, the use of the additional piece of glass after the fiber prevents this approach from being applicable in those cases when the fiber is intended to be used directly as a probe. An interesting scheme employing two fibers and six prisms was demonstrated (Clark et al., Opt. Lett. 26:1320 (2001)), in which both spectral and temporal compression were used to achieve output pulses with nearly the same duration as the input pulse. As a result of third-order dispersion, the output pulse was still distorted, but was significantly shorter than that delivered by application of a negative prechirp. An alternative approach to minimizing nonlinearity is to utilize a fiber with a larger core size and thus reduce the effective nonlinearity. However, for maintenance of a single transverse mode, the increase in core size must be accompanied by a decrease of the core-cladding refractive-index difference, which, in turn, results in a rapid increase of bend losses.
New possibilities for tailoring the optical properties of the guided modes are created with the use of recently developed microstructured fibers (MFs) (Birks et al., Electron. Lett. 31:1941 (1995) and Knight et al., Science 282:1476 (1998)). MFs can have a significantly larger effective core—cladding index difference than that of conventional fibers, and as observed (Ranka et al., Opt. Lett. 25:796 (2001)) with small core diameters, even under conditions in which the fiber is multimode, the fundamental mode can be robustly excited such that for all intents and purposes the fiber behaves as a SMF. Furthermore, the bend losses are minimal. However, it appears that there has not yet been an investigation of how large the core can become and yet still maintain this effective single-mode behavior. In addition, any such fiber must readily allow coupling from free space to the fundamental mode without significant sensitivity to input coupling conditions.
The present invention is directed to overcoming these deficiencies in the art.